The invention relates generally to the field of cardiac pacers, and in particular to cardiac pacer sub-systems including magnet rate, low battery indication, power supply, emergency modes, output formation circuits and output programming.
The physical characteristics of the human heart lend themselves to various interactive artificial pacing systems. There are two major pumping chambers in the heart, the left and right ventricles. Simultaneously contracting, these chambers expel blood into the aorta and the pulmonary artery. Blood enters the ventricles from the left and right atria, respectively. The atria are small antechambers which contract in a separate action which precedes the major ventricular contraction by an interval of about 100 milliseconds (ms), known as the AV delay. The contractions arise from a wave of electrical excitation which begins in the right atrium and spreads to the left atrium. The excitation then enters the atrio-ventricular node which delays its passage via the bundle of His into the ventricles.
Appearing in the electrocardiogram, a small signal known as the P-wave accompanies atrial contraction while a much larger signal, known as the QRS complex, with a predominant R-wave, accompanies ventricular contraction. The P and R waves can be very reliably detected as timing signals by electrical leads in contact with the respective heart chambers.
The typical implanted cardiac pacer operates by supplying missing stimulation pulses on a pacing lead attached to the ventricle. The R-wave can be sensed by the same lead. An additional lead contacts the atrium to sense P-waves, if desired. In AV sequential pacers, discussed below, the atrial lead is also used for atrial stimulation.
Cardiac pacers are useful in treating a number of cardiac disorders such as heart block caused by impairment of the ability of the bundle of His to conduct normal excitation from the atrium to the ventricle. The pacer itself is a battery powered, hermetically sealed, completely self-contained electronic device which is implanted in the body at a suitable site such as the shoulder or axillary region within an inch from the surface of the skin. The distal ends of the pacer leads are connected inside the heart to the right atrium and right ventricle and extend through a suitable blood vessel to the pacer. The proximal end of the pacer lead is taken out through an opening in the blood vessel and electrically connected to the pacer. Inside the pacer, the stimulation pulses are formed by a pulse generator. In the past, pulse generators have taken several forms but fall into two general categories: (1) those where the pulse generator consists of an R-C timing circuit and (2) those where oscillations in the output of a high frequency clock (R-C or crystal oscillator) are counted by digital circuitry. In circuits of the second kind, the pulse generator typically comprises a digital counter and logic circuitry for producing an output pulse when a given number of clock pulses is counted and means for resetting the counter in response to spontaneous or stimulated activity. An early example is found in U.S. Pat. No. 3,557,796 to Keller et al assigned to the assignee of the present application. With the miniaturization of stored program data processors, microprocessor cardiac pacer systems have given rise to more complex and yet more flexible counting arrangements. For example, a cardiac period number may be placed into a register which is regularly incremented and tested by software instructions. If the register has been counted up to the programmed number, the software branches to direct the formation of a stimulation pulse, as in "Multi-Mode Microprocessor-Based Programmable Cardiac Pacer" U.S. patent application Ser. No. 207,003, filed Nov. 14, 1980 by Leckrone et al, assigned to the assignee of the present application, and incorporated herein by reference in its entirety. In this application as in other AV sequential systems, the output circuits for the two channels, i.e. atrium and ventricle, are separate circuits.
It is extremely important to optimize the level of electrical stimulation. The charge density in the myocardium surrounding the bare electrode at the distal end of the pacer lead determines the muscular reaction. Several factors are known to affect this charge density including the amplitude of the stimulation pulse current, the voltage, the duration of the stimulation or "pulse width", the type of electrode including the area of contact and the resistance of the contacting tissue and electrochemical factors as well as the type of lead system used, i.e. unipolar or bipolar. In unipolar systems, the ground terminal is on the pacer itself while in bipolar systems the end of the pacer lead contains two spaced contacts, one of which would be regarded as ground.
Advances in pacer technology have enabled pulse parameters such as rate, width and amplitude to be altered by an externally generated programming signal, for example, using a succession of magnetic pulses to actuate a tiny reed switch in the pacer. In the past, the charge density delivered to the myocardium has been programmable by means of a variable voltage output circuit, a variable constant current output circuit, or a variable pulse width. Once a pacer is implanted and in operation at a selected pulse width and amplitude, it is extremely difficult at a later date for a physician not privy to the current parameter information, to ascertain the exact level of stimulation without knowing the amplitude beforehand. With pacers having fixed (i.e., known) amplitude and variable pulsewidth outputs, one can easily determine the applied stimulation level by gauging the pulsewidth. On the other hand, in providing for a wide range of stimulation levels in a single pacer, it has been found to be more effective to vary the amplitude. However, the stimulation level cannot then be easily determined by superficial electrical measurements.
Cardiac pacers are life-supporting, therapeutic, medical devices. They are surgically implanted and remain within a living person's body for years. The vital considerations in cardiac pacing tend to dictate a conservative approach, if not reluctance, toward commercially exploiting new developments in electronic circuitry. These tendencies are enhanced by the fact that the relatively simple functional requirements of prior art pacers have been easily implemented using pre-existing well-established hardware circuit configurations and also by the state of the art in compact batteries which limits current drain to avoid unnecessary surgical replacements. The chief objective is reliability followed closely by compactness and low current drain.
Two of the critical factors are battery depletion and failure of the main timing circuit. Many systems have been proposed for indicating low battery capacity. In particular, the magnet rate has been used for this purpose before. To observe pacer performance, the physician places a permanent magnet over the pacer which switches the pacer into an asynchronous, i.e. fixed rate, pacing mode in which the pacer's effect on the patient can be easily observed. In the prior art, a low battery indication is provided by sensing the battery voltage and by switching the fixed rate to a substantially different lower rate when a particular battery threshold is reached. See, for example, U.S. Pat. No. 4,095,603 to Davies, assigned to the assignee of the present application, the "Microlith P" Programmable Cardiac Pacer marketed by Cardiac Pacemakers, Inc., and the aforementioned copending application. In a digital pacer, main oscillator failure is difficult to cope with since all the timing for the pacer is based on the nominal frequency of the oscillator output. It is known, however, that when crystal oscillators fail, they usually fail catastrophically and quit vibrating altogether or begin vibrating at multiples of the nominal output frequency.